In nuclear emission tomography, gamma cameras or detectors typically are used for locating and displaying human glands and organs and associated abnormalities. Abnormalities may be represented by higher uptake or lower uptake than the surrounding tissue. More specifically, and with respect to using a gamma camera, gamma-ray-emitting tracer material is administered to a patient, and the tracer material is more greatly absorbed by the organ of interest than by the other tissues. The gamma camera generates data, or an image, representing the distribution of such tracer material within the patient.
A gamma camera includes a multi-channel collimator and a gamma ray detector which converts energy from the gamma ray into an electrical signal which can be interpreted to locate the position of the gamma ray interaction in the planar detector. One known gamma ray detector which is commonly used is an Anger gamma camera, which is described in H. O. Anger, "Scintillation Camera", Rev. Sci. Instrum., Vol. 29, p. 159 (1958). Another known detector is a multi-crystal scintillation detector which has an array o)f small crystals coupled to an array of light detectors, which may be either photomultipliers or photodiodes. Yet another known detector is a solid-state position sensitive detector which converts energy from the gamma ray into an electrical charge which can be detected by an array of contacts.
The Anger gamma camera includes a large scintillation crystal responsive to radiation stimuli, i.e., gamma rays emitted by the patient. An array of photomultiplier tubes typically are optically coupled to the crystal. In operation, the gamma rays emitted by the patient in the direction of the detector are collimated onto the crystal, and each gamma ray which interacts with the crystal produces multiple light events. The multiple light events are detected by photomultipliers adjacent to the point of interaction. The photomultiplier tubes, in response to the light events, produce individual electrical outputs. The signals from the array of photomultipliers are combined using analog and digital circuitry to provide an estimate of the location of the gamma ray event. Further, analog and digital processing is used to produce more accurate position coordinates to form the acquired image.
More particularly, to generate an image, a representation of the distribution of events in the crystal is generated by utilizing a matrix of storage registers whose elements are in one-to-one correspondence with elemental areas of the crystal. The crystal elemental areas are identified by coordinates. Each time a light event occurs in the crystal, the event coordinates are identified and the register in the storage register matrix corresponding to the identified event coordinates is incremented. The contents of a given register in the matrix is a number that represents the number of events that have occurred within a predetermined period of time within an elemental area of the crystal. Such number is directly proportional to the intensity of radiation emitted from an elemental area of the radiation field. The number stored in the register therefore is used to establish the brightness of a display picture element corresponding to the crystal elemental area. The distribution of a radiation field is displayed in terms of the brightness distribution of the display.
In emission tomography, a plurality of such images are taken at various view angles around the organ of interest. Typically, in transaxial tomography, a series of images, or views, are taken at equal angular increments around the patient. The series of views around the patient are reconstructed to form transaxial slices, that is, slices across the axis of rotation. The process of acquiring the views and reconstructing the transaxial slices is termed emission computed tomography (ECT) or single photon emission computed tomography (SPECT).
Most detectors used for tomography are fixed to a large bearing which allows the detector to rotate about a fixed axis (roll axis) in order to acquire the views at different angles. A particular form of gantry, known as a ring stand, allows the detector to rotate and also swivel (tilt) and pitch in order to be able to image various organs in different patient attitudes. The ring stand can be used in connection with a patient table having its long axis parallel to the roll axis, and the detector can be brought close to the patient by adjusting the pitch axis.
The collimators used in known cameras and detectors have a multiplicity of holes through which gamma rays can pass. The holes are separated by dense material, typically lead, which attenuates the gamma rays and absorbs a large fraction of the gamma rays which impinge on the dense material. The dimensions of the openings and the thickness of the lead between the holes are selected to obtain an appropriate trade-off between resolution and sensitivity as well as to minimize the penetration through the walls of the holes. Typically, the collimators are exchangeable and the operator can select a most appropriate collimator for the imaging application and the energy of the gamma rays. The array of collimator holes are typically parallel with one another, but some collimators are arranged so that the openings converge at a line or point some distance from the collimator front surface so as to obtain some magnification of the patient.
Known multi-channel collimators used in emission imaging have an image resolution which degrades linearly with an increasing distance of the object from the collimator surface. It is beneficial, therefore, to acquire each view with the collimator as close to the patient as possible. To facilitate positioning of the collimator relative to the patient, the pitch axis may be adjusted to bring the detector either closer to or farther away from the patient.
For emission tomography imaging, in addition to positioning the collimator as close to the patient as possible, it is important to precisely control the detector attitude relative to the patient. Particularly, it is important to ensure that data be acquired with the collimator holes viewing substantially transverse to the axis of rotation. Incorrect detector attitude can result in degradation of image quality.
Although controlling detector attitude is important, it is difficult to monitor the detector attitude due to changes in gantry pitch. Particularly, as the gantry is pitched to adjust for a patient's size, the detector attitude relative to the patient changes. More specifically, when the pitch axis is changed, the direction of the collimator holes will change in the longitudinal direction by the same angle as the change in the arm angle. Therefore, after changing the gantry pitch, the detector tilt must be altered to properly position the detector for data acquisition.
One known method of maintaining proper detector attitude requires a system operator to manually tilt the detector to keep the detector at a desired orientation. However, such manual adjustment is both time consuming and cumbersome, particularly if the detector is unbalanced.
It would be desirable to automatically control detector attitude so that the detector is maintained in a desired orientation as the gantry is pitched. It also would be desirable to provide such control without significantly increasing the cost of the system.